U.S. patent number 7,223,242 [Application Number 10/259,251] was granted by the patent office on 2007-05-29 for ultrasound imaging system.
This patent grant is currently assigned to Teratech Corporation. Invention is credited to Peter P. Chang, Alice M. Chiang, Xingbai He, Eric R. Kischell.
United States Patent |
7,223,242 |
He , et al. |
May 29, 2007 |
**Please see images for:
( Certificate of Correction ) ** |
Ultrasound imaging system
Abstract
The present invention is directed to an ultrasound imaging
system and method for Doppler processing of data. The ultrasonic
imaging system efficiently addresses the data computational and
processing needs of Doppler processing. Software executable
sequences in accordance with a preferred embodiment of the present
invention determines the phase shift and the auto-correlation phase
of filtered image data. In a preferred embodiment, the system of
ultrasonic imaging also includes a sequence of instructions for
Doppler processing that provides the functions for demodulation,
Gauss Match filtering, auto-correlation calculation, phase shift
calculation, frame averaging, and scan conversion implemented with
Single Instruction Multiple Data (SIMD) or Multiple Instruction
Multiple Data (MIMD) instructions. In a preferred embodiment, the
ultrasound imaging system of the present invention includes a
processing module; and memory operable coupled to the processing
module, wherein the memory stores operational instructions that
cause the processing module to map serial data to vector
representation, calculate an auto-correlation function of the data,
calculate a phase shift of the auto-correlation function to
generate a monotonic function covering all values of the phase
shift corresponding to a range of Doppler velocities and display
the resultant images, for example, as color images.
Inventors: |
He; Xingbai (Andover, MA),
Chang; Peter P. (Burlington, MA), Kischell; Eric R.
(Pepperell, MA), Chiang; Alice M. (Weston, MA) |
Assignee: |
Teratech Corporation
(Burlington, MA)
|
Family
ID: |
25511882 |
Appl.
No.: |
10/259,251 |
Filed: |
September 27, 2002 |
Prior Publication Data
|
|
|
|
Document
Identifier |
Publication Date |
|
US 20030100833 A1 |
May 29, 2003 |
|
Related U.S. Patent Documents
|
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
|
09966810 |
Sep 28, 2001 |
|
|
|
|
Current U.S.
Class: |
600/454;
600/443 |
Current CPC
Class: |
A61B
8/488 (20130101); G01S 7/52071 (20130101); G01S
7/5208 (20130101); G01S 7/52082 (20130101); G01S
7/52084 (20130101); G01S 15/8981 (20130101); A61B
8/4427 (20130101); A61B 8/54 (20130101); A61B
8/5223 (20130101); A61B 8/486 (20130101); A61B
8/463 (20130101); A61B 8/065 (20130101); A61B
8/56 (20130101); G01S 7/52044 (20130101); G01S
7/003 (20130101); G01S 7/52063 (20130101); G01S
15/584 (20130101); G01S 15/8979 (20130101) |
Current International
Class: |
A61B
8/06 (20060101) |
Field of
Search: |
;600/437,472,455,451,116,493,459,300,488,461,460,443,447,441,443.4,47,454-458
;128/916 ;73/861.25 ;700/4-5,73 ;701/18-19,48 ;708/5 ;710/69
;712/2,7,20-22 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
|
|
|
|
|
|
|
1016875 |
|
Jul 2000 |
|
EP |
|
10-73658 |
|
Mar 1998 |
|
JP |
|
WO 00/31634 |
|
Jun 2000 |
|
WO |
|
Other References
Kasai, Chihiro; et al., Real-Time-Dimensional Blood Flow Imaging
Using an Autocorrelations Technique, IEEE Transactions on Sonics
and Ultrasonics, vol. SU-32, No. 3, pp. 458-464 May 3, 1985. cited
by other.
|
Primary Examiner: Jaworski; Francis J.
Attorney, Agent or Firm: Weingarten, Schurgin, Gagnebin
& Lebovici LLP
Parent Case Text
CROSS REFERENCES TO RELATED APPLICATIONS
The present application is a continuation-in-part of U.S. patent
application Ser. No. 09/966,810, filed Sep. 28, 2001.
The entire contents of the above application is incorporated herein
by reference in its entirety.
Claims
What is claimed:
1. A ultrasound system for scanning a region of interest with
ultrasound energy to image a fluid moving within the region of
interest comprising: a transducer array; a data processing system
connected to the transducer array and including a Doppler
processing system having at least one parallel computing element,
and a memory coupled to the at least one parallel computing
element, the memory having stored therein a sequence of
instructions to vectorize radio frequency serial input data for
parallel processing, to adjust the pulse repetition frequency and
to execute color Doppler processing.
2. The ultrasound system of claim 1, wherein the at least one
parallel computing element comprises a multiplier and an adder.
3. The ultrasound system of claim 1, wherein the at least one
parallel computing element executes at least one of Single
Instruction Multiple Data instructions and Multiple Instruction
Multiple Data instructions.
4. The ultrasound system of claim 1, wherein Doppler processing is
executed in a computation device having an MMX or floating point
processor.
5. The ultrasound system of claim 1, wherein the Doppler processing
further comprises directional power Doppler, power Doppler and
pulsed wave Doppler and provides corresponding images.
6. The ultrasound system of claim 1, further comprising adjustable
controls for one of at least scan area selection, velocity display,
steering angles, color inversion, color gain, color priority, color
persistence, color baseline and frame rate.
7. In an ultrasound system for Doppler processing, a method for
imaging a region of interest comprising of: providing a transducer
array for capturing data from a region of interest; providing a
data processing system coupled to the transducer array and having a
microprocessor with at least one parallel processing element and a
memory coupled to the microprocessor having stored therein a
sequence of instructions to vectorize radio frequency serial input
data for parallel processing and execute Doppler processing
including one of color Doppler, directional power Doppler, power
Doppler and pulsed wave Doppler; Selectively adjusting the pulse
repetition frequency; and performing a Doppler processing operation
to form an image of the region of interest.
8. A computer readable medium having stored therein instructions
for causing a processing unit to execute the steps of the method of
claim 7.
9. The method for imaging a region of interest of claim 7, wherein
the sequence of instructions to vectorize the serial input data
further comprises of: mapping a stream of serial RF data to a
vector representation; determining an auto-correlation function of
the data; and determining a phase shift of the auto-correlation
function to generate a monotonic function for all values of the
phase shift corresponding to a range of Doppler velocities.
10. The method of imaging a region of interest of claim 9, wherein
the sequence of instructions further comprises: demodulating the
data to obtain in-phase and quadrature sample data; filtering the
data to remove low frequency signals; averaging a plurality of
frames of data; converting phase shift data to an index; converting
the phase shift data from scan to raster coordinates; and
displaying a plurality of images.
11. The method for imaging a region of interest of claim 7, wherein
at least one parallel processing element comprises a multiplier and
an adder.
12. The method for imaging a region of interest of claim 7, wherein
at least one parallel processing element executes at least one of
Single Instruction Multiple Data instructions and Multiple
Instruction Multiple Data instructions.
13. The method for imaging a region of interest of claim 7, wherein
the Doppler processing further comprises displaying at least one of
the presence, direction, and relative velocity of blood flow,
perfusion and contour of lumens.
14. The method of imaging a region of interest of claim 7, further
comprising providing adjustable controls for at least one of scan
area selection, velocity display, steering angles, color inversion,
color gain, color priority, color persistence, color baseline and
frame rate.
15. The method of claim 7 further comprising processing floating
point values.
16. An ultrasound imaging apparatus comprising: at least one
processing module; and memory operable coupled to the at least one
processing module, wherein the memory stores operational
instructions that cause the at least one processing module to: map
serial data to vector representation; calculate an auto-correlation
function of the data; calculate a phase shift of the
auto-correlation function to generate a monotonic function for all
values of the phase shift corresponding to a range of Doppler
velocities; convert the phase shift to an index; and display a
plurality of images including at least one of color Doppler,
directional power Doppler, power Doppler and pulsed wave
Doppler.
17. The apparatus of claim 16, wherein the apparatus further
comprises a handheld probe operable to transmit an ultrasound image
and to obtain data indicative of the flow characteristics of the
region of interest.
18. The apparatus of claim 16, wherein the plurality of images
includes color Doppler images, directional power Doppler images,
power Doppler images and pulsed wave Doppler images.
19. The apparatus of claim 16, further comprising adjustable
controls for one of at least scan area selection, velocity display,
steering angles, color inversion, color gain, color priority, color
persistence, color baseline and frame rate.
20. The apparatus of claim 16 further comprising demodulating the
data to obtain in phase and quadrature sample data.
Description
BACKGROUND OF THE INVENTION
Ultrasonic diagnostic equipment has become an indispensable tool
for clinical use. For approximately the past twenty years,
real-time B-mode ultrasound imagers are used for investigating all
soft tissue structures in the human body. One of the recent
developments within medical imaging technology is the development
of Doppler ultrasound scanners. Doppler ultrasound is an important
technique for non-invasively detecting and measuring the velocity
of moving structures, and particularly to display an estimate of
blood velocity in the body in real time.
The basis of Doppler ultrasonography is the fact that reflected
and/or scattered ultrasonic waves from a moving interface undergoes
a frequency shift. In general the magnitude and the direction of
this shift provides information regarding the motion of this
interface. How much the frequency is changed depends upon how fast
the object or moving interface is moving. Doppler ultrasound has
been used mostly to measure the rate of blood flow through the
heart and major arteries.
There are several forms of depiction of blood flow in medical
Doppler imaging or more generally different velocity estimation
systems that currently exist: Color Flow imaging, power Doppler and
Spectral sonogram. Color flow imaging (CFI), interrogates a whole
region of the body, and displays a real-time image of mean velocity
distribution. CFI provides an estimate of the mean velocity of flow
with a vessel by color coding the information and displaying it,
super positioned on a dynamic B-mode image or black and white image
of anatomic structure. In order to differentiate flow direction,
different colors are used to indicate velocity toward and away from
the transducer.
While color flow imaging displays the mean or standard deviation of
the velocity of reflectors, such as the blood cells in a given
region, power Doppler (PD) displays a measurement of the amount of
moving reflectors in the area, similarly to the B-mode image's
display of the total amount of reflectors. A power Doppler image is
an energy image in which the energy of the flow signal is
displayed. Thus, power Doppler depicts the amplitude or power of
the Doppler signals rather than the frequency shift. This allows
detection of a larger range of Doppler shifts and thus better
visualization of small vessels. These images give no velocity
information, but only show the direction of flow. In contrast,
spectral Doppler or spectral sonogram utilizes a pulsed wave system
to interrogate a single range gate or sampling volume, and displays
the velocity distribution as a function of time. The sonogram can
be combined with the B-mode image to yield a duplex image.
Typically, the top side displays a B-mode image of the region under
investigation, and the bottom displays the sonogram. Similarly, the
sonogram can also be combined with the CFI or PD image to yield a
triplex image. The time for data acquisition is then divided
between acquiring all three sets of data, and the frame rate of the
images is typically decreased, compared to either CFI or duplex
imaging.
The current ultrasound systems require extensive complex data
processing circuitry in order to perform the imaging functions
described herein. Doppler processing for providing two-dimensional
depth and Doppler information in color flow images, power Doppler
images and/or spectral sonograms require millions of operations per
second. There exists a need for an ultrasound imaging system that
provides for compute-intensive systems and methods to efficiently
address the data processing needs of information, such as Doppler
processing.
SUMMARY OF THE INVENTION
The present invention is directed to an ultrasound imaging system
and method for Doppler processing of data. The ultrasonic imaging
system efficiently addresses the data computational and processing
needs of Doppler processing. Software executable sequences in
accordance with a preferred embodiment of the present invention
determines the phase shift and the auto-correlation phase of
filtered image data. In a preferred embodiment, the system of
ultrasonic imaging also includes a sequence of instructions for
Doppler processing that provides the functions for demodulation,
Gaussian Match filtering, auto-correlation calculation, phase shift
calculation, frame averaging, and scan conversion.
In a preferred embodiment, the processing system includes parallel
processing elements which execute Single Instruction Multiple Data
(SIMD) or Multiple Instruction Multiple Data (MIMD) instructions. A
computer having a Pentium.RTM. III processor including MMX.TM.
technology is an exemplary computational device of a preferred
embodiment of the ultrasonic imaging system in accordance with the
present invention.
A method of the present invention includes imaging a region of
interest with ultrasound energy using a portable ultrasound imaging
system which in turn includes a transducer array within a handheld
probe. An interface unit is connected to the handheld probe with a
cable interface. The interface unit has a beamforming device
connected to a data processing system with another cable interface.
Output signals from the interface unit are provided to the handheld
probe to actuate the transducer array, which in turn delivers
ultrasound energy to the region of interest. The ultrasound energy
returning to the transducer array is collected from the region of
interest and transmitted from the handheld probe to the interface
unit. A beamforming operation is performed with the beamforming
device in the interface unit. The method further includes
transmitting data from the interface unit to the data processing
system such that the data processing system receives a beamformed
electronic representation of the region of interest. The data
processing system has at least one parallel processing element
integrated with a microprocessor to execute a sequence of
instructions for Doppler processing and displaying of Doppler
images.
In a preferred embodiment, the ultrasound imaging system of the
present invention includes a processing module; and memory operable
coupled to the processing module, wherein the memory stores
operational instructions that cause the processing module to map
serial data to a vector representation, demodulate the data to
obtain in-phase and quadrature sample data, calculate an
auto-correlation function of the data, calculate a phase shift of
the auto-correlation function represented as a monotonic function
in the interval corresponding to the range of Doppler velocities
according to the Nyquist criterion and expressed as a simple
mathematical function, convert the phase shift to an index and
display the images, for example, as color images.
Preferred embodiments of the present invention include the portable
ultrasound system comprising one of at least a color Doppler mode,
a directional power Doppler mode, a power Doppler mode and a pulsed
wave Doppler mode. The portable ultrasound system includes
adjustable controls for one of at least scan area selection,
velocity display, steering angles, color inversion, color gain,
color priority, color persistence, color baseline and frame
rate.
The foregoing and other features and advantages of the system and
method for ultrasound imaging will be apparent from the following
more particular description of preferred embodiments of the system
and method as illustrated in the accompanying drawings in which
like reference characters refer to the same parts throughout the
different views.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a block diagram of a preferred embodiment of the
ultrasound imaging system in accordance with the present
invention;
FIGS. 2A and 2B are flow charts of a preferred embodiment of a
method for Doppler processing in accordance with the present
invention;
FIG. 3 is a diagram illustrating a preferred embodiment of a method
of data mapping for parallel computation in accordance with the
present invention; and
FIGS. 4A and 4B are graphical representations of the Doppler phase
shift calculations in accordance with preferred embodiments of the
present invention.
FIG. 5 is a view of a display screen showing an ultrasound image
illustrating the color Doppler mode in accordance with a preferred
embodiment of the present invention.
FIGS. 6A and 6B illustrate a view of a display screen showing the
color Doppler mode of an ultrasound image and the available
controls, respectively, in accordance with a preferred embodiment
of the present invention.
FIGS. 7A-7D illustrate the controls and image sizes available by
adjusting the scan area, respectively, in accordance with a
preferred embodiment of the present invention.
FIGS. 8A-8D illustrate the controls available for adjusting
different parameters such as pulse repetition frequency, wall
filter, steering angle and color inversion for a color Doppler mode
in accordance with a preferred embodiment of the present
invention.
FIGS. 9A and 9B illustrate the color Doppler reference bar that is
used for a color invert adjustment in accordance with a preferred
embodiment of the present invention.
FIG. 10 illustrates additional adjustments available to a user in
the color Doppler mode such as color gain, color priority, color
persistence and color baseline in accordance with a preferred
embodiment of the present invention.
FIGS. 11A and 11B illustrate the images and controls provided by
direction power Doppler, respectively, in accordance with a
preferred embodiment of the present invention.
FIGS. 12A and 12B illustrate an image and the controls provided to
a user in a power Doppler mode in accordance with a preferred
embodiment of the present invention.
FIGS. 13A-13C illustrate a real-time mixed mode (B-mode scan), a
pulsed Doppler waveform and the controls provided to a user in a
pulsed wave Doppler mode in accordance with a preferred embodiment
of the present invention.
The drawings are not necessarily to scale, emphasis instead being
placed upon illustrating the principles of the invention.
DETAILED DESCRIPTION OF THE INVENTION
The ultrasound imaging system is directed at a Doppler processing
system in a portable ultrasound system. In a preferred embodiment,
the ultrasonic imaging system includes parallel computation units
and a memory having stored therein instructions to process data and
display ultrasound images using computer-efficient methods.
A preferred embodiment of the ultrasound imaging system includes a
pulse-Doppler processor for color flow imaging or map applications.
Color flow (CF) imaging combines in a single modality the abilities
of ultrasound to image tissue and to investigate blood flow. CF
images consist of Doppler information that can be color-encoded and
superimposed on a B-mode gray-scale image.
Color-flow imaging is a mean velocity estimator. There are two
known different techniques for computing the mean velocity. First,
in a pulsed Doppler system, Fast-Fourier Transforms (FFTs) can be
used to yield the velocity distribution of the region of interest,
and both the mean and the variance of the velocity profile can be
calculated and displaced as a color flow imaging. The other
approach is the one dimensional auto-correlation technique
described by Kasai et al in "Real-Time Two Dimensional Blood Flow
Imaging Using an Auto-correlation Technique" in the IEEE
Transactions on Sonics and Ultrasonics Vol. SU-32, No. 3 in May
1985, the entire contents of which are incorporated herein by
reference.
Mean blood flow velocity is estimated from the frequency spectra of
echoes. An estimate of the mean velocity in the range gate or
sample volume gives an indication of the volume flow rate. As the
frequency of the range or depth gated and sampled signal is
proportional to the velocity, the spatial mean velocity can be
determined by the mean angular frequency of P(.omega.) and is
expressed as: .PI..intg..infin..infin..times..omega..times.
.times..function..omega..times.
.times.d.omega..intg..infin..infin..times..function..omega..times.
.times.d.omega. ##EQU00001## where P(.omega.) is the power density
spectrum of the received, demodulated signal. Equation (1) gives
the mean Doppler frequency shift due to the blood flow. The mean
blood flow velocity v can then be estimated by the following
equation: .PI..omega..times. .times..times. .times..times.
.times..theta. ##EQU00002## where c is the velocity of sound and
.theta. the angle between the sound beam and the blood flow
vector.
The extent of turbulence in blood flow may be inferred from the
variance of the spectrum. Since the Doppler frequency directly
relates to the flow vector, i.e., flow direction and speed, in an
ultrasonic sample volume, the spectrum spread broadens in
accordance with flow disturbance. While in laminar flow, the
spectrum spread is narrow, since a uniform flow vector gives a
singular Doppler frequency shift. The mean angular frequency can be
determined by the phase-shift of auto-correlation of the complex
signal z(t). The inverse Fourier transform of the power density
spectrum is the auto-correlation function R(.tau.) and is expressed
as:
.function..tau..intg..infin..infin..times..function..omega..times.e.times-
. .times..omega..times. .times..tau..times.
.times.d.omega..ident..function..tau..times.e.times.
.times..PHI..function..tau. ##EQU00003## From the moment's theorem
of Fourier transforms, it can be shown that
.function..times..intg..infin..infin..times..omega..times.
.times..function..omega..times. .times.d.omega..times.
.times..function..intg..infin..infin..times..function..omega..times.
.times.d.omega. ##EQU00004## It follows then .PI..function..times.
.times..function. ##EQU00005## Therefore, the mean velocity
estimation can be reduced to an estimation of the auto-correlation
and the derivative of the auto-correlation. The estimator given by
the above expression can be calculated when data from two returned
lines are used. From Equation (3), {dot over (R)}(0)=jA(0){dot over
(.phi.)}(0) and R(0)=A(0) (7) Substituting the above equations into
(6), we have
.PI..PHI..function..apprxeq..PHI..function..PHI..function..PHI..function.
##EQU00006## Generally, .phi.(1) can be determined by either of the
following methods .PHI..function..function..times..function..times.
.times..times..function..PHI..function..function.
.times..times..function..function..PHI..function..function..times.
.times..times..function..function. ##EQU00007## In a preferred
embodiment of the ultrasonic imaging system,
.function..PHI..function..times..function..PHI..times. ##EQU00008##
represents the phase-shift, where .function.>< ##EQU00009##
.function..PHI..function..times..function..PHI..times. ##EQU00010##
is a monotonic function of .phi.(1)in the interval (-.pi., +.pi.).
Thus, every value of
.function..PHI..function..times..function..PHI..times. ##EQU00011##
uniquely defines a .phi.(1) in the interval (-.pi., +.pi.) and vice
versa. Further, sign(.phi.(1))=sign(sin(.phi.(1)))=sign(Im{R(1)})
(13)
.function..PHI..function..function..PHI..function..times..function..funct-
ion. ##EQU00012## And therefore:
.function..PHI..function..times..function..PHI..function..times..function-
..times..function..function..times..function..function.
##EQU00013## Similarly, the .phi.(1) can be determined based on
.function..PHI..function..times.
.times..PHI..function..function..PHI..function..times..function..function-
..times..function..function. ##EQU00014##
In a preferred embodiment of the present invention, the estimator
given by the above expressions 15 and/or 15(a) can be calculated
when data from at least two returned lines are used. The equations
15 and 15a represent the phase shift .phi. as a monotonic function
in the interval between negative and positive pi (-.pi.,+.pi.) and
express the phase shift as simple mathematical calculations. In a
preferred embodiment of the present invention, more lines are used
in order to improve the signal-to-noise ratio. Data from several RF
lines are needed in order to get useful velocity estimates by the
auto-correlation estimator. Preferably, in a particular embodiment
between eight (8) and sixteen (16) lines are acquired for the same
image direction. The lines are divided into range gates throughout
the image depths, and the velocity is estimated along the
lines.
The CFI pulses are interspersed between the B-mode image pulses in
duplex imaging. It is known that a longer duration pulse train
gives an estimator with a lower variation. However, a good spatial
resolution necessitates a short pulse. In a particular embodiment
of the ultrasound imaging system of the present invention, a
separate pulse is preferably used for the B-mode image, because the
CFI pulse is too long for high quality gray-scale image.
While Color Flow Imaging (CFI) sonograph has been an effective
diagnostic tool in clinical cardio-vascular application, power
Doppler (PD) imaging provides an alternative method of displaying
the blood stream in the insonified regions of interest. While CF
imaging displays the mean or standard deviation of the velocity of
reflectors such as, for example, blood cells in a given region, PD
displays a measurement of the amount of moving reflectors in the
area, similarly to the B-mode image's display of the total amount
of reflectors. Thus, power Doppler is akin to a B-mode image with
the stationary reflectors suppressed. This is particularly useful
for viewing small moving particles with small scattering
cross-sections such as red blood cells.
The power Doppler image displays the integrated Doppler power
instead of the mean frequency shift as used for color Doppler
imaging. As discussed hereinbefore, the color flow mapping is a
mean frequency estimation, and can be expressed as:
.omega..intg..infin..infin..times..omega..times.
.times..function..omega..times.
.times.d.omega..intg..infin..infin..times..function..omega..times.
.times.d.omega. ##EQU00015## where .omega. represents the mean
frequency shift and P(.omega.) is the power density spectrum of the
received signal. It has also been shown hereinbefore by inverse
Fourier transform that
.function..tau..intg..infin..infin..times..function..omega..times.e.-
times. .times..omega..times. .times..tau..times. .times.d.omega.
##EQU00016## The total integrated Doppler power can be expressed as
.intg..infin..infin..times..function..omega..times.d.omega.
##EQU00017## Substituting Equation (16) into (17), it follows that
.function..intg..infin..infin..times..function..omega..times..function..t-
imes. .times..omega..times.
.times..times.d.omega..intg..function..omega..times.d.omega.
##EQU00018## and that the 0th lag of the auto-correlation function
can be used to compute the integrated total Doppler power. In other
words the integrated power in the frequency domain is the same as
the integrated power in the time domain and hence the power Doppler
can be computed in one preferred embodiment from either time-domain
or the frequency-domain data. In either embodiment, the unwanted
signal from the surrounding tissue such as the vessel walls is
removed via filtering.
In a preferred embodiment, the PD computation can be carried out in
software executing a sequence of instructions on the host processor
similarly to the computation of the CFI processing described
hereinbefore. In a preferred embodiment, parallel computation units
such as those in the Intel.RTM. Pentium.RTM. and Pentium.RTM. III's
MMX.TM. coprocessors that allow rapid computation of the required
functions are used. Intel.RTM.'s MMX.TM. is a Pentium.RTM.
microprocessor that executes applications faster than non-MMX.TM.
Pentium.RTM. microprocessor. It is designed to improve the
performance of multimedia and communication algorithms. The
technology includes new instructions and data types which achieve
higher levels of performance for these algorithms or host
processors. In particular, MMX.TM. Pentium.RTM. microprocessors
have microprocessor instructions that are designed to handle video,
audio and graphical data more efficiently. Further, MMX.TM.
technology consists of a Single Instruction Multiple Data (SIMD)
process which makes it possible for one instruction to perform the
same operation on multiple data items. In addition, the memory
cache on the MMX.TM. Processor has increased to, for example, 32
thousand bytes, which provides for fewer accesses to memory that is
off the microprocessor. In an alternate preferred embodiment, a
Digital Signal Processor (DSP) can also be used to perform the
processing function. Such architecture permits flexibility in
changing digital signal processing algorithms and transmitting
signals to achieve the best performance as the region of interest
is changed.
As discussed hereinbefore, the frequency content of the Doppler
signal corresponds to the velocity distribution of the blood. It is
common to device a system for estimating blood movement at a fixed
depth in tissue. A transmitter emits an ultrasound pulse that
propagates into and interacts with tissue and blood. The
backscattered signal is received by the same transducer and
amplified. For a multiple-pulsed system, one sample is acquired for
each pulse emitted. A display of the distribution of velocities can
be made by Fourier transforming the received signal and displaying
the result. This display is also called a sonogram. Often a B-mode
image is presented along with the sonogram in a duplex system, and
the area of investigation or range gate is displayed on the image.
The placement and size of the range gate are determined by the
user, and this determines the time instance for the sampling
operation. The range gate length determines the area of
investigation and sets the length of the emitted pulse.
The calculated spectral density is displayed on a screen with
frequency on the y-axis and time on the x-axis. The intensity at a
point on the screen indicates the amplitude of the spectrum and is,
thus, proportional to the number of blood scatterer moving at a
particular velocity.
FIG. 1 is a schematic functional block of one embodiment of the
ultrasound imaging system 10 of the invention. Similar imaging
systems are described in U.S. Pat. No. 5,957,846 to Alice M. Chiang
et al., issued Sep. 28, 1999, entitled "Portable Ultrasound Imaging
System," the entire contents of which are being incorporated herein
by reference. As shown, the system 10 includes an ultrasonic
transducer array 14 which transmits ultrasonic signals into a
region of interest or image target 12, such as a region of human
tissue, and receives reflected ultrasonic signals returning from
the image target. The system 10 also includes a front-end interface
or processing unit 18 which is connected by cables 16, for example,
coaxial cables to the transducer array 14 and includes a transducer
transmit/receive control chip 22.
Ultrasonic echoes reflected by the image target 12 are detected by
the ultrasonic transducers in the array 14. Each transducer
converts the received ultrasonic signal into a representative
electrical signal which is forwarded to an integrated chip having
preamplification circuits and time-varying gain control (TGC)
circuitry 30. The preamplification circuitry sets the level of the
electrical signals from the transducer array 14 at a level suitable
for subsequent processing, and the TGC circuitry is used to
compensate for attenuation of the sound pulse as it penetrates
through human tissue and also drives the beamforming circuits 32 to
produce a line image. The conditioned electrical signals are
forwarded to the beamforming circuitry 32 which introduces
appropriate differential delay into each of the received signals to
dynamically focus the signals such that an accurate image can be
created. Further details of the beamforming circuitry 32 and the
delay circuits used to introduce differential delay into received
signals and the pulses generated by a pulse synchronizer are
described in U.S. Pat. No. 6,111,816 to Alice M. Chiang et al.,
issued Aug. 29, 2000 entitled "Multi-Dimensional Beamforming
Device," the entire content of which are being incorporated herein
by reference.
A memory 30 stores data from a controller 28. The memory 30
provides stored data to the transmit/receive chip 22, the TGC 30
and the beamformer 32. The output from the system controller 28 is
connected directly to a custom or FireWire Chipset. The FireWire
Chipset is described in co-pending U.S. patent application Ser. No.
09/449,780, entitled "Ultrasound Probe with Integrated
Electronics," by Jeffrey M. Gilbert et al., the entire contents of
which are being incorporated herein by reference. "FireWire" refers
to IEEE standard 1394, which provides high-speed data transmission
over a serial link. There also exists a wireless version of the
FireWire standard allowing communication via an optical link for
untethered operation.
The FireWire standard and an ultrasound probe with integrated
electronics as described in co-pending U.S. patent application Ser.
No. 09/791,491, entitled "Ultrasound Probe With Integrated
Electronics," by Alice M. Chiang et al., the entire contents of
which are being incorporated herein by reference, may be used in
preferred embodiments of the present invention. The FireWire
standard is used for multimedia equipment and allows 100-200 Mbps
and preferably in the range of 400-800 Mbps operation over an
inexpensive 6 wire cable. Power is also provided on two of the six
wires so that the FireWire cable is the only necessary electrical
connection to the probe head. A power source such as a battery or
IEEE1394 hub can be used. The FireWire protocol provides both
isochronous communication for transferring high-rate, low-latency
video data as well as asynchronous, reliable communication that can
be used for configuration and control of the peripherals as well as
obtaining status information from them. Several chipsets are
available to interface custom systems to the FireWire bus.
Additionally, PCI-to-FireWire chipsets and boards are currently
available to complete the other end of the head-to-host connection.
CardBus-to-FireWire boards can also be used.
Although the VRAM controller directly controls the ultrasound scan
head, higher level control, initialization, and data processing and
display comes from a general purpose host such as a desktop PC,
laptop, or palmtop computer. The display can include a touchscreen
capability. The host writes the VRAM data via the VRAM Controller.
This is performed both at initialization as well as whenever any
parameters change (such as number or positions of zones, or types
of scan head) requiring a different scanning pattern. During
routine operation when data is just being continually read from the
scan head with the same scanning parameters, the host need not
write to the VRAM. Because the VRAM controller also tracks where in
the scan pattern it is, it can perform the packetization to mark
frame boundaries in the data that goes back to the host. The
control of additional functions such as power-down modes and
querying of buttons or dial on the head can also be performed via
the FireWire connection.
Although FireWire chipsets manage electrical and low-level protocol
interface to the FireWire interface, the system controller has to
manage the interface to the FireWire chipset as well as handling
higher level FireWire protocol issues such as decoding asynchronous
packets and keeping frames from spanning isochronous packet
boundaries.
Asynchronous data transfer occurs at anytime and is asynchronous
with respect to the image data. Asynchronous data transfers take
the form of a write or read request from one node to another. The
writes and the reads are to a specific range of locations in the
target node's address space. The address space can be 48 bits. The
individual asynchronous packet lengths are limited to 1024 bytes
for 200 Mbps operation. Both reads and writes are supported by the
system controller. Asynchronous writes are used to allow the host
to modify the VRAM data as well as a control word in the controller
which can alter the operation mode. Asynchronous reads are used to
query a configuration ROM (in the system controller FPGA) and can
also be used to query external registers or I/O such as a "pause"
button. The configuration ROMs contain a querible "unique ID" which
can be used to differentiate the probe heads as well as allow
node-lockings of certain software features based on a key.
Using isochronous transfers, a node reserves a specified amount of
bandwidth and it gets guaranteed low-overhead bursts of link access
every 1/8000 second. All image data from the head to the host is
sent via isochronous packets. The FireWire protocol allows for some
packet-level synchronization and additional synchronization is
built into the system controller.
The front-end processing or interface unit system controller 28
interfaces with a host computer 20, such as a desktop PC, laptop or
palmtop, via the custom or FireWire Chipsets 24, 34. This interface
allows the host to write control data into the memory 26 and
receive data back. This may be performed at initialization and
whenever a change in parameters such as, for example, number and/or
position of zones, is required when the user selects a different
scanning pattern. The front-end system controller 28 also provides
buffering and flow control functions, as data from the beamformer
is sent to the host via a bandwidth-constrained link, to prevent
data loss.
The host computer 20 includes a keyboard/mouse controller 38, and a
display controller 42 which interfaces with a display or recording
device 44. A graphical user interface described in co-pending U.S.
patent application Ser. No. 09/822,764 entitled "Unitary Operator
Control for Ultrasonic Imaging Graphical User Interface," by
Michael Brodsky, the entire contents of which are being
incorporated herein by reference, may be used in a preferred
embodiment of the present invention.
The host computer further includes a processing unit such as
microprocessor 36. In a preferred embodiment of the ultrasound
imaging system in accordance with the present invention the
microprocessor 36 includes on-chip parallel processing elements. In
a preferred embodiment, the parallel processing elements may
include a multiplier and an adder. In another preferred embodiment,
the processing elements may include computing components, memories,
logic and control circuits. Depending on the complexity of the
design, the parallel processing elements can execute either SIMD or
Multiple Instruction Multiple Data (MIMD) instructions.
Further, the host computer includes a memory unit 40 that is
connected to the microprocessor 36 and has a sequence of
instructions stored therein to cause the microprocessor 36 to
provide the functions of down conversion, scan conversion, M-mode,
and Doppler processing which includes color flow imaging, power
Doppler and spectral Doppler, and any post-signal processing. The
down conversion or mixing of sampled analog data may be
accomplished by first multiplying the sampled data by a complex
value and then filtering the data to reject images that have been
mixed to nearby frequencies. The outputs of this down-conversion
processing are available for subsequent display or Doppler
processing.
The scan conversion function converts the digitized signal data
from the beamforming circuitry 32 from polar coordinates
(r,.theta.) to rectangular coordinates (x,y). After the conversion,
the rectangular coordinate data can be forwarded for optional post
signal processing where it is formatted for display on the display
44 or for compression in a video compression circuit. Scan
conversion and beamforming and associated interfaces are described
in U.S. Pat. No. 6,248,073 to Jeffrey M. Gilbert et al., issued on
Jun. 19, 2001, entitled "Ultrasound Scan Conversion with Spatial
Dithering," the entire contents of which are being incorporated
herein by reference.
The Doppler processing (CFI, PD, spectral Doppler) is used to image
target tissue 12 such as flowing blood. In a preferred embodiment,
with pulsed Doppler processing, a color flow map is generated. In a
preferred embodiment, the CFI, PD, Spectral Doppler computation can
be carried out in software running on the host processor. Parallel
computation units such as those in the Intel.RTM. Pentium.RTM. and
Pentium.RTM. III's MMX.TM. coprocessors allow rapid computation of
the required functions. For parallel processing computation, a
plurality of microprocessors are linked together and are able to
work on different parts of a computation simultaneously. In another
preferred embodiment, digital Signal Processor (DSP) can also be
used to perform the task. Such arrangement permits flexibility in
changing digital signal processing algorithms and transmitting
signals to achieve the best performance as region of interest is
changed.
Single Instruction Multiple Data (SIMD) parallel processors allow
one micro-instruction to operate at the same time on multiple data
items to accelerate software processing and thus performance. One
chip provides central coordination in the SIMD parallel processing
computer. Currently, SIMD allows the packing of four single
precision 32-bit floating point values into a 128-bit register.
These new data registers enable the processing of data elements in
parallel. Because each register can hold more than one data
element, the processor can process more than one data element
simultaneously. In a preferred embodiment of the present invention,
all the data is organized efficiently to use SIMD operations. In a
particular embodiment, Multiple Instruction Multiple Data (MIMD)
parallel processors may be used, which include a plurality of
processors. Each processor can run different parts of the same
executable instruction set and execute these instructions on
different data. This particular embodiment employing MIMD may be
more flexible than the embodiment utilizing SIMD, however may be
more expensive. All the kernel functions such as demodulation,
Gauss match filtering, Butterworth high pass filtering,
auto-correlation calculation, phase-shift calculation, frame
averaging, color-averaging, spatial domain low-pass filtering, and
scan conversion interpolation are implemented with SIMD or MIMD.
The Doppler processing results in the processed data being scan
converted wherein the polar coordinates of the data are translated
to rectangular coordinates suitable for display or video
compression.
The control circuit, preferably in the form of a microprocessor 36
inside of a personal computer (e.g., desktop, laptop, palmtop),
controls the high-level operation of the ultrasound imaging system
10. The microprocessor 36 or a DSP initializes delay and scan
conversion memory. The control circuit 36 controls the differential
delays introduced in the beamforming circuitry 32 via the memory
26.
The microprocessor 36 also controls the memory 40 which stores
data. It is understood that the memory 40 can be a single memory or
can be multiple memory circuits. The microprocessor 36 also
interfaces with the post signal processing functional instructions
and the display controller 44 to control their individual
functions. The display controller 44 may compress data to permit
transmission of the image data to remote stations for display and
analysis via a transmission channel. The transmission channel can
be a modem or wireless cellular communication channel or other
known communication method.
The preferred embodiments of the ultrasonic imaging system address
the main problem of performing Doppler processing by software which
typically is the speed for processing in real time. FIG. 2
illustrates a preferred embodiment of a method 100 for Doppler
processing in accordance with the present invention.
The method 100 includes mapping or vectorizing of the serial input
RF data for parallel computation per step 102. Further details of
the data mapping process are described in FIG. 3 which
diagrammatically illustrates a preferred embodiment of a method 150
for data mapping for parallel processing. An input data steam 152
is mapped into an input vector representation. The vectors 154, 156
are sequentially provided to the microprocessor's 36 parallel
processing elements in which an instruction can be executed that
allows all data within the vectors 154, 156 to be operated on in
parallel. The dimension of the vector P is determined by the
hardware constraints of the microprocessor and may equal the number
of parallel processors. For example, current MMX.TM. technology
allows a vector with four dimensions, thus limiting the processing
of four dimension vectors in parallel. The dimensions of the vector
representation are not limited to the current available technology
and can accommodate increases in the dimension of the vector
representation.
The method 100 includes the step 104 of smoothing the RF data to
enhance the signal to noise ratio (SNR), for example, by curve
fitting techniques. The RF data is then demodulated to generated IQ
data, where I represents in-phase and Q represents quadrature
samples in step 106. Demodulation may be performed, but is not
limited to, by a rectangular filter.
In a preferred embodiment, Doppler processing is operated on each
sub-segment instead of each single sample. This method saves
processing time because the number of sub-segments is much less
than number of samples. This is important for clinical use of
software-based Doppler products, where a real-time system is
critical. Further, by processing sub-segments the sensitivity to
flows increases. The data processed per a sequence of instructions
in a preferred embodiment represents the average over a segment,
which has higher signal to noise ratio.
The use of data segments may however cause lower spatial
resolution. To reduce such an effect, the signal sequence is not
divided as individual segments, instead each adjoining segment is
overlapped with each other. In other words, the center distance or
down-sample space between each segment is less than the length of
segments. According to the Nyquist sampling theorem, if such a
distance is less or equal to the half segment length, the spatial
resolution can be recovered by proper interpolation methods. For
example, if there exists a demodulated IQ data sequence y.sub.0,
y.sub.1, y.sub.2, y.sub.3, y.sub.4, y.sub.5, y.sub.6, y.sub.7, the
segment length as four and down-sample space as two, a down-sampled
data sequence z.sub.0, z.sub.1, Z.sub.2, is generated where z.sub.0
is the average of y.sub.0, y.sub.1, y.sub.2, y.sub.3, z.sub.1 is
average of y.sub.2, y.sub.3, y.sub.4, y.sub.5, z.sub.2 is the
average of y.sub.4, y.sub.5, y.sub.6, y.sub.7. Mathematically, the
average over each segment can be expressed as
.function..times..times. .times..function..function. ##EQU00019##
where y(t) is demodulated complex IQ data, .times.
.times..function..ltoreq..ltoreq. ##EQU00020## If the down-sample
space is .DELTA.t.ltoreq.L/2, then the down-sampled data can be
expressed as z.sub.s(k)=z(k.DELTA.t) where k=0,1,2, (21)
The method 100 further includes the step 108 of filtering the
down-samples using, for example, without limitation, the Gauss
Match filter. A match filter is used to maximize the
signal-to-noise ratio. The signal for Doppler processing is
generated by performing quadrature demodulation with the emitted
frequency (.omega..sub.0) and then Gauss Match filtering the
complex signal. y(t)=Gauss(t)*[rf(t)exp(j .omega..sub.0t)] (22)
where rf(t) is the raw RF data. To save the calculation time, match
filtering is only performed on down-sampled samples. From (19), it
follows z(t)=Re ct(t)*Gauss(t)*[rf(t)exp(j .omega..sub.0t)] (23)
Equation (23) can be also expressed as z(t)=Gauss(t)*x(t) (24)
x(t)=Re ct(t)*[rf(t)exp(j .omega..sub.0t)] (25) Thus, the
down-sampled data can be expressed as .function..function..times.
.times..DELTA..times.
.times..intg..infin..infin..times..function..tau..times..function..times.
.times..DELTA..times. .times..tau..times. .times.d.tau.
##EQU00021##
The method 100 further includes the step 110 of calculating the
power along the sample line before a high pass filter or wall
filter. Per step 112, stationary signals are filtered out by a high
pass filter, for example, but not limited to, a Butterworth filter.
The method 100 then proceeds to step 114 wherein the
auto-correlation function R(1) of the signal is calculated.
Per step 116, the phase shift of the auto-correlation phase is then
calculated efficiently. As described hereinbefore, the mean
velocity can be determined by the mean angular frequency
.PI..intg..infin..infin..times..PI..times.
.times..function..PI..times.
.times.d.PI..intg..infin..infin..times..function..PI..times.
.times.d.PI. ##EQU00022## where P( .omega.) is the power density
spectrum of Doppler signal. The mean blood flow velocity .nu. can
then be estimated by the following equation .PI..omega..times.
.times..times. .times..times. .times..theta. ##EQU00023## where c
is the velocity of sound and .theta. the angle between the sound
beam and the blood flow vector. It has been shown that:
.PI..PHI..function..apprxeq..PHI..function..PHI..function..PHI.
##EQU00024## Generally, .phi.(1) can be determined by either of the
following methods
.PHI..function..function..times..function..times..function..PHI..function-
..function..times..function..function..PHI..function..times..function..fun-
ction. ##EQU00025## All these operations are too time consuming to
be implemented by software. Besides, these functions are not
monotonic in the interval between -pi and +pi (-.pi., +.pi.). Thus,
not all values of the phase shift corresponding to the range of
Doppler velocities according to the Nyquist criterion are accounted
for. In a preferred embodiment of the Doppler processing system in
accordance with the present invention
.function..PHI..function..times..function..PHI..times. ##EQU00026##
represents the phase-shift, where .function.>< ##EQU00027##
.function..PHI..function..times..function..PHI..times. ##EQU00028##
is a monotonic function of .phi.(1) in the interval (-.pi.,+.pi.).
In other words, every value of
.function..PHI..function..times..function..PHI..times. ##EQU00029##
uniquely defines a .phi.(1) in the interval (-.pi., +.pi.) and vice
versa. Calculation of the sin.sup.2 function avoids the use of a
square root operation, which is computationally intensive. As we
know
.function..PHI..function..function..PHI..function..times..function..funct-
ion..times.
.times..function..PHI..function..times..function..PHI..function..times.
.times..function..times..function..function..times..function..function.
##EQU00030## Similarly, the angle phi .phi.(1) may be also
calculated by .function..PHI..function..times. .times..PHI..times.
.times..function..PHI..function..times..function..function..times..functi-
on..function. ##EQU00031##
The method 100 further includes the step 118 wherein the echo
signals are rejected by color priority. In a preferred embodiment,
if the power before the high pass filter is larger than a
predetermined threshold, the phase shift is set to zero. Per step
120 the phase shift is smoothed by, but is not limited to, a
spatial low-pass filter. The phase-shift is averaged with previous
frames such as, for example, the previous two frames per step 122.
The phase shift is converted to a color index per step 124 to
obtain a curve like the one represented by the phase shift
calculation in equations 35 and 35a. In a preferred embodiment, a
look up table may be used, but is not limited to, in order to
convert the phase shift to a color index. Color averaging is then
performed per step 126. Per step 128 the phase shift represented in
scan coordinates is transformed or converted to raster coordinates
using, for example, but is not limited to, interpolation methods.
The resultant color images are superpositioned on the B-mode image
in step 130.
FIG. 4A illustrates a graphical representation 170 of the Doppler
phase shift calculations, in particular for the following equations
10 and 15 described hereinbefore. The graphical representation of
equation 15 in accordance with a preferred embodiment, is a
monotonic function in the interval between -pi and pi which spans
the range of Doppler velocities according due to the Nyquist
criterion. This can be expressed by simple mathematical operations,
as shown in the right hand side of equation 15.
.PHI..function..function..times..function..function..function..PHI..funct-
ion..times..function..PHI..function..times..times.
.times..times..function..function. ##EQU00032##
FIG. 4B illustrates a graphical representation 190 of equations 9
and 15a.
.PHI..function..function..times..function..times..function..function-
..PHI..function..times. .times..PHI..times.
.times..function..PHI..times.
.times..times..function..function..times..function..function.
##EQU00033## Equation 15a, is the graphical representation of a
preferred embodiment, and is a monotonic function in the interval
between -pi and pi which spans the range of Doppler velocities
according to the Nyquist criterion, and can be expressed by simple
mathematical operations, as shown in the right hand side of
equation 15a.
It should be noted that an operating environment for the system 10
includes a processing system with at least one high speed
processing unit and a memory system. In accordance with the
practices of persons skilled in the art of computer programming,
the present invention is described with reference to acts and
symbolic representations of operations or instructions that are
performed by the processing system, unless indicated otherwise.
Such acts and operations or instructions are sometimes referred to
as being "computer-executed", or "processing unit executed."
It will be appreciated that the acts and symbolically represented
operations or instructions include the manipulation of electrical
signals by the processing unit. An electrical system with data bits
causes a resulting transformation or reduction of the electrical
signal representation, and the maintenance of data bits at memory
locations in the memory system to thereby reconfigure or otherwise
alter the processing unit's operation, as well as other processing
of signals. The memory locations where data bits are maintained are
physical locations that have particular electrical, magnetic,
optical, or organic properties corresponding to the data bits.
The data bits may also be maintained on a computer readable medium
including magnetic disks, optical disks, organic disks, and any
other volatile or non-volatile mass storage system readable by the
processing unit. The computer readable medium includes cooperating
or interconnected computer readable media, which exist exclusively
on the processing system or is distributed among multiple
interconnected processing systems that may be local or remote to
the processing system.
In preferred embodiments, the ultrasound imaging system includes
color and power Doppler modes to detect the presence, direction and
relative velocity of blood flow by assigning color-coded
information to these parameters. The color is depicted in a region
of interest that is overlaid on to the B-mode image. All forms of
ultrasound-based imaging of red blood cells are derived from the
echo signal that is received in response to the transmitted signal.
The primary characteristics of this signal are its frequency and
its amplitude (or power). The frequency shift is determined by the
movement of the red blood cells. The amplitude is dependent on the
amount of blood in motion present within the volume that is sampled
by the ultrasound beam.
In a preferred embodiment, as illustrated in FIG. 5, color Doppler
allows flow velocity information to be detected over a portion of
the B-Mode image. The mean Doppler shift is then displayed against
a gray scale representation of the structures. Color Doppler images
display blood flow by mapping the frequency shift (velocity) of
blood cells. Flow towards the transducer is assigned shades of red,
and flow away from the transducer is displayed in shades of blue.
Alternate embodiments can use different color contrasts. Higher
frequencies are displayed in lighter colors and lower frequencies
in darker colors. For example, the proximal carotid artery is
normally displayed in hues of bright red and orange because the
flow is toward the transducer and the frequency (velocity) of flow
in this artery is relatively high. By comparison, the flow in the
jugular vein is displayed as blue because it flows away from the
transducer.
The color Doppler mode provides, in preferred embodiments, fourteen
examination types to provide presets depending upon the probes
used. The range of WF includes 1% to 25% of pulse repetition
frequency (PRF), and steering angles include, for example,
+/-15.degree., +/-20.degree., and 0 for flat linear probes.
Further, maps in red/blue based on velocity of blood flow and high
spatial resolution versus high frame rate for certain probes is
also provided. In a preferred embodiment, as illustrated in FIG. 6A
and 6B the color Doppler (CD)-mode for image display is selected by
choosing the CD-mode button from a image mode bar, or from a Modes
menu. To select the CD image control setting, the user selects the
CD tab 242 at the bottom of the image control bar. The scan area
controls the size of the region of interest. Further, in preferred
embodiments the Doppler systems use a wall filter to reduce and
preferably eliminate unwanted low frequency signals from the
display. The color controls 244a-d can be adjusted to increase the
quality of an image.
FIGS. 7A-7D illustrate the controls and the image sizes available
in a color Doppler mode in accordance with a preferred embodiment.
The size of the scan area is one of the controls the user
manipulates to affect the frame rate. The smaller the scan area,
the faster the frame rate. Alternatively, the larger the scan area,
the slower the frame rate. For example, for cardiac or arterial
applications, a small scan area can be used to accurately visualize
the flow dynamics. A medium or large scan can also be used for
applications where the blood flow dynamics do not change rapidly
over time, or if the user wants to get a larger overall view of the
blood flow.
In a preferred embodiment, as shown in FIG. 8A, the ultrasound
system provides for adjustment of the pulse repetition frequency
(PRF). The PRF adjusts the range of the velocity display shown on a
color bar. Decreasing the PRF values improves the display of slow
blood flow. The maximum value for the PRF is dependent on the
specific probe that is being used and how deep the region of
interest is. The PRF can be set high enough to prevent aliasing,
and low enough to provide adequate detection of low flow. It may be
necessary to vary the PRF during an examination, depending on the
speed of the blood flow, and/or if a pathology is present. If the
PRF is set too high, low frequency shifts caused by low velocity
flow may not be shown. In preferred embodiments, the PRF is set
higher for cardiac and arterial applications than it is for venous
or small parts application.
In a preferred embodiment, Doppler systems use a wall filter to
eliminate unwanted low frequency signals from the display. Raising
the wall filter reduces the display of low velocity tissue motion.
As illustrated in FIG. 8B, the ultrasound system provides for an
adjustment of a wall filter. As the wall filter is lowered, more
information is displayed, however, more wall tissue motion can also
be displayed. The wall filter can range from 1% to 25% of the PRF
(in increments of 2, 10, 25, 50 or 100 depending on the wall filter
value). The wall filter is set high enough so that color Doppler
flash artifacts from tissue or wall motion are not displayed, but
low enough to display slow flow. If the wall filter is set too
high, slower flow may not be seen. The wall filter is set higher
for applications where there is significant tissue motion (such as
a cardiac application), or in instances where the probe is moved
rapidly while color Doppler is on. It may be set lower for small
parts or instances where flow is slow but there is not much tissue
motion.
A preferred embodiment also provides for the adjustment of a
steering angle as shown in FIG. 8C. The steering angle is used only
on flat linear arrays, and allows adjustments for optimal angle
between beam and flow. The angle range is, but not limited to, for
example, 0, +/-15.degree., +/-20.degree.. Doppler angle adjustment
is important in obtaining adequate signals and in interpreting the
color Doppler examination. Changes in the angle of the Doppler beam
to the path of flow causes changes in the color Doppler display.
When using color Doppler, the user has to be aware of the Doppler
angle to flow. At a 90 degree angle to flow, an absent or confusing
color pattern is displayed (even when the flow is normal). An
adequate Doppler angle to flow is required in order to obtain
useful color Doppler information. In most instances, the lower the
angle of the color Doppler beam to flow, the better the received
signal. Electronic steering is useful in the embodiments where the
flow is at a poor angle to the color Doppler beam. However, in many
embodiments it is also necessary to use the "heel to toe"
technique, which involves using pressure on one end of the probe or
the other to improve the Doppler angle to flow. It should be noted
that curved linear and phased array probes do not have the
capability of electronic steering and depending on the clinical
situation, may require the use of the "heel and toe" technique to
obtain an adequate angle to flow. As color Doppler is angle
dependent, a weak flow signal results if the vessel is
approximately at 90 degrees to the beam.
Another preferred embodiment includes the provision of color
inversion. When the user chooses the Color Invert box illustrated
in FIG. 8D, the color Doppler reference bar is inverted as well as
the color within the region of interest. The color reference bar is
divided by a zero baseline. To invert the color Doppler reference
bar, the user can click on the Color Invert box. The Play button is
used to see the Color Invert effect on a frozen image.
Conventionally, the color red is assigned to positive frequency
shifts (flow toward the probe), and blue is assigned to negative
frequency shifts (or flow away from the probe). However, this color
assignment can be reversed at the user's discretion by selecting
Invert. Whether or not the user has inverted the display, flow
toward the probe is assigned the colors of the top half of the
color bar, and flow away from the probe is assigned the colors of
the bottom half of the color bar. FIGS. 9A and 9B illustrate the
inverted and non-inverted reference bar in the color Doppler mode.
Invert may be used when scanning the internal carotid artery (ICA),
for example. In general, flow in this vessel goes away from the
probe. If Invert is enabled, the ICA flow is displayed in shades of
red. The color bar then displays shades of blue on the top half,
and a shade of red on the bottom.
Preferred embodiments also allow for the adjustment of the color
gain which increases or decreases the amplification of the
returning signal displayed or being played. The color gain
adjustment 352 is illustrated in FIG. 10. The color gain can be
increased whenever the fill of color within a vessel is inadequate,
and can be decreased whenever an unacceptable amount of color is
seen outside of a vessel.
Preferred embodiments also accommodate the adjusting of the color
priority of the image which defines the amount of color displayed
over bright echoes and helps confine color within the vessel walls.
This adjustment in color priority 354 is illustrated in FIG. 10.
This control affects the level at which color information
overwrites the B-mode information. If the user needs to see more
flow in an area of some significant B-mode brightness, they
increase the color priority. If a user wants to better contain the
display of flow within the vessel(s), they decrease the color
priority. If, however, the color priority is set all the way to the
left, no color is displayed.
A preferred embodiment also accommodates for adjusting the color
persistence as illustrated in FIG. 10. The color persistence
adjustment 356 averages frames of color. Increasing the persistence
causes the display of flow to persist on the 2D image. Decreasing
the persistence allows better detection of short duration jets, and
provides a basis for better flow/no flow decisions. Adjusting color
persistence also produces better vessel contour depiction. To
adjust the color persistence for a scan, in a preferred embodiment,
the slider is moved to the right or left to achieve the desired
image. When color persistence is set high, the saved image (single
frame) may not look exactly the same as when the image is saved.
The user receives a warning when this occurs; however, this does
not occur when exporting images.
In a preferred embodiment, as illustrated in FIG. 10, the color
baseline 358 can also be adjusted. The baseline refers to the zero
baseline within the color Doppler window. Adjusting this control
moves the zero baseline up or down. To adjust the baseline, the
slider is moved to the right or left to achieve the desired range
of degree. In general, it is not necessary to adjust the color
baseline. If the user does adjust it, moving the baseline down
displays more positive flow and moving the baseline up displays
more negative flow in an embodiment. When the user adjusts the
color baseline, they are able to display more forward or reversed
flow. This adjustment can be used to prevent aliasing in either
direction.
Further embodiments include an adjustment for high frame rate 360
shown in FIG. 10. This control adjusts the line density within the
Doppler region of interest. The user can choose between an image
with higher line density resulting in better spatial resolution or
lower line density resulting in a better frame rate. The high frame
rate is used when the flow rate is high such as in cardiac or
certain arterial applications. This control is available with
certain embodiments of the probes. When the user selects high
spatial resolution, line density is increased. However, when the
user selects high frame rate, line density is decreased.
Another preferred embodiment of the present invention includes a
directional power Doppler mode which can be viewed as a combination
of both conventional power Doppler and color Doppler modes. It
provides the same increased sensitivity as conventional power
Doppler as well as directional information derived from color
Doppler. Directional power Doppler does not provide an estimate of
the frequency (velocity) of blood flow. The color palette is
proportional to the strength of the Doppler signal. The mode allows
the user to achieve good image quality of deep arteries and other
tissue. One also has the option to apply a high frame rate or high
resolution to control the quality of the scan. The benefits of
directional power Doppler include examination types to provide
presets (dependent upon certain probes), Doppler steering angles of
(+/-15.degree.,=/-20.degree., 0.degree., for linear probes),
without limitation, range of WF and increments (2 to PRF/4 Hz) and
increased Doppler sensitivity performance.
As illustrated in FIGS. 11A and 11B, the DirPwr-mode 382 is
selected and for image display, the DirPwr-mode button from the
image mode bar can be chosen, or from the Modes menu. To select the
DirPwr image control setting, the user selects the DirPwr tab at
the bottom of the image control bar. Similar to the color Doppler
mode, the scan area controls-the size of the region of interest.
The Doppler systems use the wall filter to eliminate unwanted low
frequency and provides color control adjustments.
As illustrated hereinbefore, the scan can be adjusted even in the
directional power Doppler mode. The size of the scan area is one of
the controls that is used to affect the frame rate. The smaller the
scan area, the faster the frame rate. Alternatively, the larger the
scan areas, the slower the frame rate. For cardiac or arterial
applications, a small scan area is used to accurately visualize the
flow dynamics. A medium or large scan can also be used for
applications where the blood flow dynamics do not change rapidly
over time, or if the user wishes to get a larger overall view of
the blood flow. The region of interest can be defined by adjusting
the scan area. The scan area options range from small, medium to
large.
In the directional power Doppler mode, the PRF can be set high
enough to prevent aliasing, and low enough to provide adequate
detection of low flow. It may be necessary to vary the PRF during
one examination depending on the speed of the blood flow, and/or if
the pathology is present. If the PRF is set too high, signals
caused by slow flow states with few red blood cells may not be
shown. In an embodiment, the PRF is set higher for high flow states
than it is for low flow states. The PRF can be set high enough to
prevent aliasing, and low enough to provide adequate detection of
low flow. It may be necessary to vary the PRF during one
examination, depending on the speed of the blood flow, and/or if
the pathology is present. If flow is weak or slow it is displayed
in darker shades. If flow is strong or fast it is shown in brighter
shades. Directional power Doppler is somewhat angle dependent. Good
Doppler angles can be maintained.
Further, Doppler systems use a wall filter to eliminate unwanted
low frequency signals from the display. Raising the wall filter
reduces the display of low velocity tissue motion. As the wall
filter is lowered, more information is displayed, however, more
wall tissue motion is also displayed. The wall filter can be set
high enough so that Doppler flash artifacts from tissue or wall
motion are not displayed, but can be set low enough to display slow
flow. If the wall filter is set too high, slower flow may not be
seen. The wall filter is set higher for applications where there is
significant tissue motion, or in instances where the probe is moved
rapidly while directional power Doppler is enabled. It may be set
lower for small parts or instances where flow is weak but there is
not much tissue motion. The wall filter range is dependent on which
probes the user is using as well as the PRF setting. Wall filter
range is from 1% to 25% of the PRF (in increments of 2, 10, 25, 50
or 100 depending on the wall filter value).
In addition, the steering angle adjusts the optimal angle to flow
in the direction power Doppler mode. The steering angle is used
only on linear arrays, and allows adjustments for optimal angle
between beam and flow. The angle range is without limitation, 0,
+/-15.degree., +/-20.degree.. The steering angle adjusts the
optimal angle to flow. Doppler angle adjustment is important in
obtaining adequate signals and in interpreting the color Doppler
examination. Changes in the angle of the Doppler beam to the path
of flow causes changes in the color Doppler display. When using
color Doppler the user has to be aware of the Doppler angle to
flow. At a 90 degree angle to flow, an absent or confusing color
pattern is displayed (even when the flow is normal). An adequate
Doppler angle to flow is required in order to obtain useful color
Doppler information. In most instances, the lower the angle of the
color Doppler beam to flow, the better the received signal.
Electronic steering is available for flat linear array probes.
Electronic steering is useful in preferred embodiments in those
instances where the flow is at a poor angle to the color Doppler
beam. However, in many instances it is also necessary to use the
"heel and tow" technique, which involves using pressure on one end
of the probe or the other to improve the Doppler angle to flow.
Curved linear and phased array probes do not have the capability of
electronic steering and depending on the clinical situation, may
require you to use the "heel and toe" technique to obtain an
adequate angle to flow. To change the steering angle, the down
arrow is clicked and the desired setting is selected.
In an embodiment including directional power Doppler, the user can
adjust the color gain which increases or decreases the
amplification of the returning signal displayed or being played.
Color gain can be increased whenever the fill or color within a
vessel is inadequate, and should be decreased whenever an
unacceptable amount of color is seen outside of a vessel.
Further, in an embodiment including directional power Doppler,
color priority can be adjusted. This control affects the level at
which color information overwrites the B-mode information. If the
user needs to see more flow in an area of some significant B-mode
brightness, the color priority is increased. If the user wants to
better contain the display of flow within the vessel(s), they
decrease the color priority. To adjust the color priority the
slider is moved to the right or left to achieve the desired range.
Raising the priority displays color on brighter structures.
Lowering priority increases containment of the display of flow to
within the vessels. If, however, color priority is set all the way
to the left, no color will be displayed.
The color persistence can be adjusted which averages frames of
color. Increasing the persistence causes the display of flow to
persist on the 2D image. Decreasing the persistence allows better
detection of short duration jets, and provides a basis for better
flow/no flow decisions. Adjusting color persistence also produces
better vessel contour depiction.
Further, in the preferred embodiment it is not necessary to adjust
the color baseline. If the user does not wish to adjust it, moving
the baseline down displays more positive flow and moving the
baseline up displays more negative flow. When you adjust the color
baseline, you are able to display more forward or reversed flow.
This adjustment can be used to prevent aliasing in either one
direction.
The preferred embodiment includes a high frame rate adjustment.
This control adjusts the line density within the Doppler region of
interest. The user can choose between one image with higher line
density resulting in better spatial resolution or lower line
density resulting in better frame rate. High frame rate is used
when the flow rate is high such as in cardiac or certain arterial
applications. This control is only available on certain probes.
Conventional power Doppler in accordance with a preferred
embodiment images blood flow by displaying the density of red blood
cells, as opposed to their velocity. Large amplitude signals are
assigned a bright hue and weak signals are assigned a dim hue. All
flow is displayed in shades of the same color. No directional
information is provided in this embodiment.
In accordance with another embodiment, the ultrasound imaging
system includes a power Doppler mode. The sensitivity of power
Doppler is greater than color Doppler. Amplitude estimation in this
mode is less noisy than a mean frequency estimate, therefore more
real signal is detected and displayed with the power Doppler mode.
Power Doppler is also less angle dependent than Color Doppler
because of this increase in sensitivity, and does not alias. All
flow is displayed in shades of the same color, however, no
directional information is provided. The benefits of power Doppler
include it being more sensitive to low flow than color or direction
power Doppler, it is the preferred mode to show perfusion and
contour of vessel lumen and large amplitude signals are assigned a
bright hue and weak signals are assigned a dim hue. For example,
the jugular vein is shown in brighter colors than the carotid
artery because the vein contains more red blood cells at any given
time than does the artery.
To select the Pwr-mode for image display, the user select a
Pwr-mode button from the image mode bar, or from the Modes menu as
illustrated in FIGS. 12A and 12B. To select the Pwr image control
setting, the user selects the Pwr tab 412 at the bottom of the
image control bar. Increasing the PRF increases the range of the
display. Lower PRF values improve the display of slow flow. PRF is
dependent on the specific probe that is being used.
In the embodiment including power Doppler mode, the size of the
scan area is one of the controls the user has access to that most
affects the frame. The smaller the scan area, the faster the frame
rate. Alternatively, the larger the scan area, the slower the frame
rate. For cardiac or arterial applications, a small scan area is
used to accurately visualize the flow dynamics. A medium or large
scan can also be used for applications where the blood flow
dynamics do not change rapidly over time, or if the user wishes to
get a larger overall view of the blood flow. The region of interest
can be defined by adjusting the scan area. The scan area options
range from small, medium to large as previously described.
In the power Doppler mode, the PRF (Pulse Repetition Frequency) can
be low enough to provide adequate detection of low amplitude flow.
It may be necessary to vary the PRF during one examination
depending on the speed of the blood flow, and/or if the pathology
is present. If the PRF is set too high, signals caused by slow flow
states with few red blood cells may not be shown. The PRF is set
higher for high flow states than it is for flow states in most
preferred embodiments. To increase the PRF, the right arrow is
selected. To decrease the PRF, the left arrow in a control is
selected. The PRF values range from 100 Hz to 15 kHz and depend
upon the probe being used. If there are few red blood cells being
sampled, flow is shown in the darker shades of gold. If there are
many red blood cells being sampled, flow is displayed in a brighter
shade of gold. Power Doppler is the least angle dependent Doppler
mode. However, it is good practice to maintain reasonable Doppler
angles while using this mode.
In an embodiment including power Doppler mode, a wall filter can be
set high enough so that power Doppler flash artifacts from tissue
or wall motion are not displayed. However, the wall filter can be
set low enough to display low amplitude signals. If the wall filter
is set too high, low amplitude signals may not be seen. The wall
filter is set higher for applications where there is significant
tissue motion, or in instances where the probe is moved rapidly
while power Doppler is on. It may be set lower for small parts or
instances where the amount of blood flow is small and there is not
much tissue motion.
With respect to adjusting the steering angle, power Doppler is less
angle dependent than Color Doppler, but many of the same rules
still apply. Changes in the angle of the Doppler beam to the path
of flow causes changes in the power Doppler display. At 90 degrees
to flow, a weak signal may be displayed even when the flow is
normal. The steering angle adjusts the optimal angle to flow. To
increase the steering angle, the down arrow to the desired setting
is selected. The angle range is, for example, -0, +/-15, +/-20
without limitation. If an adequate angle is used, the Doppler angle
and the quality of the received signal is enhanced. The more red
cells being sampled the stronger the received signal. Further,
attenuation plays a role in the strength of the displayed signal.
The closer the area of interest is to the probe, the less the
attenuation and the stronger the received signal. In a preferred
embodiment, electronic steering is available for flat linear array
probes. Electronic steering is useful in those instances where the
flow is at a poor angle to the power Doppler beam. In many
instances, it is also necessary to use the "heel and toe"
technique, which involves using pressure on one end of the probe or
the other to improve the Doppler angle to flow. Curved linear and
phased array probes do not have the capability of electronic
steering and depending upon the clinical situation, may require you
to use the "heel and toe" technique to obtain an adequate angle to
flow.
Embodiments including power Doppler also provide for adjusting
color gain. Color gain can be increased whenever the fill of color
within a vessel is inadequate, and should be decreased whenever an
unacceptable amount of color is seen outside of a vessel. The user
can adjust the color gain which modifies the degree of sensitivity
to receive signals and increases or decreases the amplification of
the returning signal displayed/played. To adjust the color gain,
the slider is moved to the right or left to achieve the desired
gain. The color gain maximum amplitude is, for example, 1 to 30
units.
In a preferred embodiment including a power Doppler mode, the color
priority control affects the level at which color information
overwrites the B-mode information. If the user needs to see more
flow in an area of some significant B-mode brightness, the color
priority is increased. If the user wants to better contain the
display of flow within the vessel(s), they decrease the color
priority. To adjust the color priority the slider is moved to the
right or left to achieve the desired range. Raising the priority
displays color on brighter structures. Lowering priority increases
containment of the display of flow to within the vessels. The color
priority maximum amplitude is, for example, in the range of 1 to 30
units.
Preferred embodiments of the power Doppler mode include adjusting
the color persistence which averages frames of color. Increasing
the persistence causes the display of flow to persist on the
two-dimensional image. Decreasing the persistence allows better
detection of short duration jets, and provides a basis for better
flow/no flow decisions. Adjusting color persistence also produces
better vessel contour depiction.
In an embodiment, the user has the option of selecting a high frame
rate versus a high quality image. If the user wants a high quality
image, the button for high spatial resolution is selected. If the
user wants a high frame rate, the button for high frame rate is
selected. When the user selects high spatial resolution, line
density is increased. When the user selects high frame rate, line
density is decreased.
A preferred embodiment includes a pulsed wave Doppler mode, which
is a mode of scanning in which a series of pulses are used to study
the motion of blood flow at a small region along a desired
scanline, called the sample volume or sample gate. As illustrated
in FIG. 13A and 13B the x axis of the graph represents time while
the y axis represents Doppler frequency shift. The real-time mixed
mode is provided that shows the B-mode scan area (FIG. 13A) along
with a bottom window (FIG. 13B) that displays the pulsed Doppler
waveform. The shift in frequency between successive ultrasound
pulses caused mainly by moving red blood cells, can be converted
into velocity if an appropriate angle between the insonating beam
and blood flow is known. The strength of the signal appears as
shades of gray, for example, in this spectral display. The
thickness of the spectral signal is indicative of laminar or
turbulent flow (laminar flow typically shows a narrow band of blood
flow information). In a preferred embodiment, pulsed wave Doppler
and B-mode are shown together in a mixed mode display. This
provides the user with the capability of monitoring the exact
location of the sample volume on the B-mode image while doing
pulsed wave Doppler. The benefits of pulsed wave Doppler include a
plurality of examination types, for example, fourteen examination
types to provide presets, the range of wall filter is between 1% to
25% of the Pulsed Repetition Frequency (PRF), measurements are
available for all examination types including peak systole (PS),
end diastole (ED), Resistive Index (RI), and systole-diastole ratio
(PS/ED) and duplex imaging is available for real-time display using
B-mode and pulsed wave Doppler. The pulse wave Doppler (PWD) mode
is selected for image display by selecting a PWD mode button from
the image mode bar, or from the modes menu. The PWD image control
setting is selected using the PWD tab 452 at the bottom of the
image control bar.
In a preferred embodiment of a pulsed wave Doppler mode, there are
three sweep speeds available. Further, adjustments to the velocity
display are available which changes the unit of the vertical scale
within the pulsed wave Doppler window between cm/s and KHz. The
centimeter unit is available when the correction angle is situated
between 0 and +/-70.degree. in a preferred embodiment.
Preferred embodiments of the pulsed wave Doppler window include
adjustments to the PRF (Pulse Repetition Frequency) which adjusts
the velocity range of the display. The maximum value for PRF is
dependent on the specific probe that is being used and the location
of the sample volume. As PRF increases, the maximum Doppler shift
that can be displayed without aliasing also increases. Lower PRF
values improves the display of slow blood flow. Increasing the PRF
also increases the thermal index value.
In a preferred embodiment including the pulsed wave Doppler mode, a
wall filter (high pass frequency filter) is used to reduce and
preferably eliminate unwanted low frequency high-intensity signals
(also known as clutter) from the display. Clutter can be caused by
tissue motion or by rapid movement of the probe. A wall filter that
is high enough to remove clutter is used but that is low enough to
display spectral information near the baseline. The range of wall
filter values is, for example, between 1% and 25% of the PRF
value.
Preferred embodiment of the pulsed wave Doppler mode include
adjusting the steering angle. The steering angle is used only on
flat linear arrays, and allows adjustment for optimal angle between
beam and flow. The Doppler signal is weak when the angle between
the insonating beam and blood flow approximates 90 degrees. Angles
of 60.degree. or less are recommended for steering angle. The
steering angle does not directly affect the calibration of the
velocity scale.
A preferred embodiment also provides for adjusting an invert
option. The user can invert the pulsed Doppler waveform. The
Doppler scale is separated by a zero baseline that extends across
the width of the spectral display. The data above the baseline is
classified as forward flow. The data below the baseline is
classified as reverse flow. When inverted, the reverse flow
displays above the baseline and the forward flow appears below the
baseline. To invert the flow, an Invert box is selected.
A preferred embodiment also includes the provision of adjusting a
correcting angle in the pulsed wave Doppler mode. To obtain
accurate velocities, the user maintains Doppler angles of 600 or
less. However, the user can employ higher values of correction
angle, especially in peripheral vascular applications where the
blood vessels are more parallel to the face of the probe. In a
preferred embodiment the maximum value for the correction angle is
+/-70.degree.. Velocity display in cm/s is shown only in the range
between 0 and +/-70.degree.. Above 70.degree., the error in the
velocity calculation is too large and the velocity scale is
converted automatically to frequency, independent of the correction
angle. However, the flow direction cursor is shown on the screen
for reference. The correction angle control is active also on
frozen images. To adjust the correction angle, the user selects the
left and right arrows in the Correction Angle field and clicks to
the desired value setting. Or, in the alternative, the user can
employ the quick adjustment key to easily set the angle correction
between +/-60.degree. degrees and 0 degrees.
Further, a preferred embodiment provides for adjusting sample
volume size and position in the pulsed wave Doppler mode. The user
can adjust the size of the pulsed wave Doppler region being
examined by setting the sample volume size control. The lower the
value, the narrower the sample size used in the calculation of flow
velocity contents. To adjust the sample volume size, the user
selects the left and right arrows next to the SV Size field to set
to the desired value. The value range for sample volume size is,
for example, 0.5 to 20 mm (in 0.5 mm increment). The adjustment is
visible on the SVgate and Probe Info interface when enabled. The
position of the sample volume can be adjusted by using the
touch-pad control. Left-click on the sample volume (the line
becomes green,) is selected to move it to the desired location, and
left-click is selected again to anchor it to make active.
Alternatively, this movement can be accomplished by using the
keyboard arrow keys. Modifying the depth location of the sample
volume affects the thermal index value.
Preferred embodiments of the pulsed wave Doppler mode include the
adjustments for gain baseline and sound volume. The user can adjust
the gain which increases or decreases the amplification of the
returning signal displayed/played. The gain can be adjusted so that
the spectral waveform is bright but not so high that the systolic
window fills in, or other artifacts are created. The baseline
refers to the zero baseline within the Pulsed Wave Doppler window.
Adjusting this control moves the zero baseline up or down. When the
user adjusts the baseline, they are able to display more forward or
reversed flow, taking advantage of the full scale available at that
particular PRF value. This adjustment is visible on the reference
bar. Adjusting the sound volume control lets the user define the
volume of the pulsed wave Doppler. The sound volume of the spectral
signal can be adjusted to a comfortable level. If it is too high,
system noise may interfere with the sound produced by the blood
flow.
It should be understood that the programs, processes, methods and
systems described herein are not related or limited to any
particular type of computer or network system (hardware or
software), unless indicated otherwise. Various types of general
purpose or specialized computer systems may be used with or perform
operations in accordance with the teachings described herein.
In view of the wide variety of embodiments to which the principles
of the present invention can be applied, it should be understood
that the illustrated embodiments are exemplary only, and should not
be taken as limiting the scope of the present invention. For
example, the steps of the flow diagrams may be taken in sequences
other than those described, and more or fewer elements may be used
in the block diagrams. While various elements of the preferred
embodiments have been described as being implemented in software,
in other embodiments hardware or firmware implementations may
alternatively be used, and vice-versa.
It will be apparent to those of ordinary skill in the art that
methods involved in the system and method for ultrasound imaging
may be embodied in a computer program product that includes a
computer usable medium. For example, such a computer usable medium
can include a readable memory device, such as, a hard drive device,
a CD-ROM, a DVD-ROM, or a computer diskette, having computer
readable program code segments stored thereon. The computer
readable medium can also include a communications or transmission
medium, such as, a bus or a communications link, either optical,
wired, or wireless having program code segments carried thereon as
digital or analog data signals.
The claims should not be read as limited to the described order or
elements unless stated to that effect. Therefore, all embodiments
that come within the scope and spirit of the following claims and
equivalents thereto are claimed as the invention.
* * * * *